1. Field of the Invention
The invention relates generally to the field of imaging ultrasonic medical transducer assemblies, and, specifically, to an apparatus and method for cooling the transducer.
2. Description of the Related Art
Ultrasonic medical transducers are used to observe the internal organs of a patient. The ultrasonic range is described essentially by its lower limit: 20 kHz, roughly the highest frequency a human can hear. The medical transducers emit ultrasonic pulses which echo (i.e., reflect), refract, or are absorbed by structures in the body. The reflected echoes are received by the transducer and these received signals are translated into images. Such translation is possible because the reflections from the internal organs vary in intensity according to the “acoustic impedance” between adjacent structures. The acoustic impedance of a tissue is related to its density; the greater the difference in acoustic impedance between two adjacent tissues the more reflective their boundary will be.
The frequency of the ultrasonic beams has an effect on both the image resolution and the penetration ability of the ultrasonic device. Higher frequency ultrasound waves have a longer near field (i.e., the region in the sound beam's path where the beam diameter decreases as the distance from the transducer increases) and less divergence in the far field (i.e., the region in the sound beam's path where the beam diameter increases as the distance from the transducer increases): higher frequency ultrasonic waves thus permit greater resolution of small structures. However, high frequency ultrasonic waves have less penetrating ability because their energy is absorbed and scattered by soft tissues. On the other hand, lower frequency ultrasonic waves have a greater depth of penetration, but the received images are much less well defined. The conventional frequency range for imaging human internal organs (using sound waves) is typically from about 3 MHz to about 5 MHz.
Two types of resolution generally apply to ultrasound imaging transducers: lateral resolution and axial resolution. Lateral resolution is the ability to resolve objects side by side and, as discussed above, is proportionally affected by the frequency (the higher the frequency, the higher the lateral resolution). Higher frequency transducers are used for infants and children because there is less need for deep penetration and the smaller structures can be viewed with greater lateral resolution. Lower frequencies are used for adults where the internal structures are larger and there is a greater need for depth penetration. Of course, when determining the appropriate frequency to be used, the structure, tissue, or organ to be viewed (and the exact purpose of the imaging) can matter more than the age of the subject. For example, diagnostic breast imaging on an adult may require a frequency of about 7 MHz or higher.
Axial resolution is the ability to resolve objects that lie one above the other. Because this is related to depth penetration, axial resolution is inversely proportional to the frequency of the transducer (depending on the size of the patient). In large patients, higher frequency beams are rapidly absorbed by the objects closest to the transducer, thus reducing depth penetration and axial resolution.
The focusing of an ultrasonic transducer can be implemented in one of two ways: mechanical or electronic. Mechanical focusing consists of placing an acoustic lens on the surface of the transducer or using a transducer with a concave face. One or several piezoelectric elements are used. In order to create a sweeping beam for 2D imaging, a single element may be oscillated back and forth, several elements may be rotated, or a single element may be used with a set of acoustic mirrors. This last transducer type (with the acoustic mirrors) is sometimes called the “wobbler” because of the vibration created as the mirrors rotate or oscillate inside the housing.
Electronic focusing uses a process called phased array, where multiple piezoelectric elements in an array are stimulated (or “fired”) sequentially in order to form and focus the beam. In an annular array, circular or ringlike elements and/or arrays are used. In a linear array, a row of elements is used to form and focus the beam. A transducer contains an array of transmitting elements and a similar array of receiving elements. An example of how a linear array forms and focuses a sound beam is shown in FIG. 1. In order to focus at point X, the outer elements 101 and 107 fire first, then elements 102 and 106, then elements 103 and 105, and finally element 104. As shown in FIG. 1. the resulting wavefronts combine to form a semicircular ultrasonic pulse whose focal point is X. By varying the sequential pattern of firing, the distance of focal point X from the transducer can be changed. Furthermore, varying the sequential pattern of firing can also be used to steer the beam. Steering is used to move focal point X left and right in FIG. 1. By rapidly steering a series of beams from left to right, a 2D cross-sectional image may be formed.
In 2D mode, one sweep from left to right is a frame, and the number of sweeps in a second is the frame rate (or fps—frames per second). Conventional frame rates ranges from about 12 fps to about 30 fps. The number of beams formed over time is the Pulse Repetition Frequency (PRF), measured in pulses per second. The range of PRFs for most commercial echocardiographs is between about 200 and about 5000 pulses per second. PRF varies with the type of imaging being performed. Most of the time spent in each second is used wailing for the echoes to return to the receiving elements in the transducer. In other words, after a beam is formed, the transmitting elements lie dormant while the beam travels to the various objects and then some of that sound energy returns (as echoes) to the transducer's receiving elements. The amount of time that the transmitting elements are transmitting sound energy is called the duty factor. Most transducers are acting as a receiver about 99% of the time, in which case the duty factor is 1 (%).
Aperture is the size of the active transmitting and receiving portion of a transducer array. Aperture is measured in square centimeters and is a function of the number of transducer elements used simultaneously to form an image. A common measurement of aperture size is F-number or F#, which is defined as the ratio of depth to aperture. These values are related to the lateral resolution (LR) by the following function:   LR  =            λ      *      F      ⁢      #        =          λ      *              D        A            where λ=wavelength of sound pulse                D=depth of the scan        A=aperture of the scan        
As can be seen from the above equation, for a fixed frequency, the aperture size must increase as the scanning depth increases in order to maintain uniform lateral resolution throughout the image. Many ultrasonic systems select a transmit aperture based on the scan depth setting and continuously vary the reception aperture. It is desirable to achieve low F#s, which, because the scanning depth is limited by the position of the desired subject, is identical to seeking larger aperture sizes. It is also desirable to seek small wavelengths, which is equivalent to seeking higher frequencies.
There are a number of modes in which an ultrasonic transducer operates. The basic modes are A Mode, B Mode, M Mode, and 2D Mode. The A Mode is amplitude mode, where signals are displayed as spikes that are dependent on the amplitude of the returning sound energy. The B Mode is brightness mode, where the signals are displayed as various points whose brightness depends on the amplitude of the returning sound energy. The M Mode is motion mode, where B Mode is applied and a strip chart recorder allows visualization of the structures as a function of depth and time. The 2D Mode is two-dimensional (imaging) mode, where B Mode is spatially applied by sweeping the beam (as described above) so that structures are seen as a function of depth and width.
2D Mode refers to the most basic, fundamental imaging mode. There are other imaging modes, which also image in two dimensions (in three dimensions in some new technologies), but these are referred to by their own names, usually based on the type of technology/methodology used to produce the image. Some of these other imaging modes will be described below. When the term 2D Mode is used, it only refers to the basic spatially oriented B Mode, and not all two dimensional imaging modes.
There have been various solutions to the problem of higher frequency-greater resolution-less depth penetration and lower frequency-lower resolution-greater depth penetration. One solution is harmonic imaging. With conventional imaging, the ultrasound system transmits and receives a sound pulse of a specific frequency (the “fundamental” frequency). As discussed above, parts of the sound pulse are reflected back to the transducer, where the reflected sound at the fundamental frequency is processed. In harmonic imaging, the transducer does not listen for the fundamental frequency, but for other frequencies—most notably, the “harmonic” frequency, which is twice the fundamental frequency (this is sometimes referred to as the “first harmonic”). To be more technically accurate, the transducer receives a signal comprised of a number of frequencies, including the fundamental and harmonic, and separates out the signal at the harmonic frequency (the “harmonic” for short). Once separated out, the harmonic is processed to produce an image.
Harmonics are generated by the object being imaged, either with or without the assistance of human intervention. When assisted, harmonics ale generated by ultrasound “contrast agents” which are injected into the patient's body. These contrast agents generally contain very small bubbles which generate two kinds of echoes when struck by a pulse. First, the conventional echo is bounced back based on the fundamental frequency. But then the bubble vibrates (in response to the pulse), thereby generating a harmonic signal. Because these contrast agents are formed for the purpose of creating harmonic echoes (rather like a bell when struck by a clapper), very strong echoes are generated at the harmonic frequency, thereby producing excellent high contrast images.
When not relying on contrast agents, harmonic imaging uses the harmonics that are generated by the tissue of the body itself (having been “rung” by the ultrasonic pulse). Obviously, these harmonics do not generate as clear a picture as the harmonics from contrast agents. The ability to create harmonics in tissue varies depending on the tissue's location in the ultrasound beam's field of view. The most pronounced imaging effect is from harmonics located at and around the focal point of the ultrasonic beam (i.e., mid-field). No harmonics are generated by tissue in the near field, and the harmonics generated in the far field attenuate quickly after being produced.
Harmonic imaging has a number of advantages. The beam formed at the harmonic frequency is narrower and has lower side-lobes, thereby significantly improving grayscale contrast resolution. Furthermore, since the harmonics are generated inside the body, they only pass through the fat layer once, rather than twice.
Some other modes of imaging are dependent on the Doppler effect, the phenomena whereby the frequency of sound from an approaching object has a higher frequency and, conversely, sound from a receding object has a lower frequency. In ultrasonic systems, this effect is used to determine the velocity and direction of blood flow in a subject. Doppler techniques can also be used wraith ultrasonic transducers which operate in continuous wave mode (i.e., part of the transducer array transmits while another part simultaneously receives).
Pulsed wave Doppler effect techniques have proven to be very accurate in blood flow studies. However, it the velocity of the blood how being measured exceeds the Nyquist Limit (half the PRF), the ultrasonic readings become inaccurate. Most Doppler techniques try to achieve a high a PRF as possible in order to avoid this effect. One type of imaging, Color Flow Imaging or CFI, uses this effect (called “aliasing”) to detect flow disturbances, e.g., transitions from laminar to turbulent flow. In CFI, multiple sample volumes are detected and displayed utilizing color mapping for direction and velocity flow data. Common mapping formats are to BART (Blue Away, Red Towards), RABT (Red Away, Blue Towards), or enhanced/variance flow maps where color saturations indicate turbulence/acceleration and color intensities indicate higher velocities. Some maps use a third color, green, to indicate accelerating velocities and turbulence.
It is desirable for the ultrasonic system to operate at the highest frequency (for the reasons discussed above) and at the maximum acoustic intensity. Maximizing the acoustic intensity increases imaging performance by increasing the depth penetration and maximizing the signal to noise ratio (SNR). However, higher frequencies and greater acoustic intensities cause the ultrasonic transducer to heat up, and there are regulatory limits (and practical limits) on the surface temperature of an ultrasonic probe when interacting with a subject. Specifically, the upper temperature limit on the patient contact surface of an ultrasonic transducer is generally considered to be about 41° C. or about 16° C. above ambient temperature.
The heat of the transducer surface is generated both by the electroacoustic energy conversion process taking place in the transducer's piezoelements and by the acoustic energy passing through and/or into adjacent transducer materials (and the patient herself). Different methods and systems halve been developed to deal with the heating problem and they can be broken into two types: active and passive. Passive solutions use passive cooling mechanisms, i.e., spreading out the dissipated heat to as large an external transducer surface area as possible. Typically, the heat generated by the transducer array is absorbed by solid thermal conductors, and then this captured heat is moved by thermal convection into the transducer's external case, where it can dissipate in the atmosphere. Ideally, the external heat-convecting surface area would consist of the transducer's entire external surface area.
One example of a passive heat dissipation system is U.S. Pat. No. 5,213,103 ('103 patent), which is hereby incorporated by reference in its entirety. FIG. 2 shows the outside of the transducer in the '103 patent. A heat sink device (internal to the transducer, thus not shown) placed inside the transducer 10 behind the piezoelectric elements in the face 14 (i.e., the patient contact surface) on the head 12 of transducer 10. The heat sink extends the entire length of the transducer and conducts heat away from face 14, through head 12, to the sides of handle 16 and power cable 18. Heat conductive epoxy is used both to attach the heat sink to the transducer housing and to conduct the heat from the heat sink to the transducer housing.
Another example of a passive heat dissipation system is U.S. Pat. No. 5,555,887 ('887 patent), which is hereby incorporated by reference in its entirety. The '887 patent applies heat dissipation to an endoscopic ultrasound transducer by embedding aluminum foil in acoustic lens material in front of the transducer array. Heat is conducted by the aluminum foil to a heat sink positioned at a distance from the patient contacting surface of the probe. U.S. Pat. No. 5,721,463 ('463 patent), which is hereby incorporated by reference in its entirety, describes a passive heat dissipation system which uses a bundle of coaxial cables to vent heat away from the face of the probe.
These passive heat sinks ale effective, but they also add to the transducer's overall thermal dissipation resistance. The fundamental limitation is that, for most transducers, even if heat is spread uniformly on the external case surfaces, it only takes a few watts of transducer driving power to cause the average transducer surface temperature to become unacceptable either with respect to the patient or the sonographer. In these cases, and particularly for small transducers having small surface areas, one may find that one is unable to operate at the allowable acoustic intensity limit because of excessive temperatures.
Active solutions, on the other hand, use active cooling means, such as circulating coolant systems. One example, U.S. Pat. No. 5,560,362 ('362 patent), which is hereby incorporated by reference in its entirety, describes a heat dissipation system in which a pumping or pressurization means actively circulates a gaseous or liquid coolant in a cable, part of which is nearby the transducer array. The system can be a single pass, multipass, or closed loop circulating system, and the coolant may pass through heat exchanger, a heat pipe, a thermoelectric cooler, an evaporator/condenser system, and/or a phase change material.
An ultrasonic transducer cooling system which uses feedback control is shown in U.S. Pat. No. 6,210,356 ('356 patent), which is hereby incorporated by reference in its entirety. The '356 patent is directed to a catheter which provides ultrasonic energy (and perhaps medicine) as a therapeutic treatment to a site inside it patient's body. Thus, no imaging or sensing is being performed by the ultrasonic transducer in the '356 patent. Temperature sensors are positioned in the surface coating of the catheter next to the ultrasound transducer in order to provide a measure of the temperature on the exterior surface of the catheter. This measure is used as a feedback control signal for the power circuits of the ultrasonic transducer. After the user sets a predetermined temperature, the power circuits decrease or increase power in the same proportion as the measured temperature is above or below the predetermined temperature.
The device described in the '356 patent also includes safety control logic which detects when the temperature at a temperature sensor has exceeded a safety threshold. When this occurs, the power circuits stop providing power to the ultrasonic transducer. However, such a feedback control system can be inappropriate for ultrasonic imaging/measuring applications.
Although abruptly turning off the power during a therapeutic ultrasonic session may not be damaging, abruptly turning off the power during an imaging/measuring session can be potentially dangerous (e.g., a sudden blackout during a surgical procedure). Even when not dangerous, turning off the image makes the diagnosis and analysis of image data difficult. Thus, there is a need for a system and method for reducing the temperature of the ultrasonic transducer in general, and the patient contacting surface in specific, which does not merely turn the transducer off. Furthermore, there is a need for a system and method which may work either as a replacement for the conventional active or passive heal dissipation systems or as an adjunct to them.